The disclosure herein relates generally to ultrasound delivery apparatus (e.g., delivery systems) and ultrasound therapy provided therewith. More particularly, the disclosure herein pertains to ultrasound therapy methods, which can be used in vitro and in vivo, and ultrasound delivery apparatus for providing such ultrasound therapy. More particularly, the disclosure herein pertains to ultrasound therapy methods and systems that use ultrasound imaging therewith, e.g., ultrasound therapy applications (e.g., thermal and non-thermal therapy relating to vasculature (e.g., decrease plaque growth), nerve structure (e.g., denervation), tumor (e.g., tissue ablation or lesion formation), cardiac tissue (e.g., cardiac ablation), drug delivery (e.g., activation of drug in tissue), etc.).
Current technology for providing therapy using ultrasonic energy is inadequate. In addition, current technology for imaging based on ultrasonic signals is inadequate. High-intensity focused ultrasound (HIFU) continues to receive increased attention as a therapeutic tool in the treatment of cancer and other tissue abnormalities (see, Wu, et al., “Advanced hepatocellular carcinoma: Treatment with high-intensity focused ultrasound ablation combined with transcatheter arterial embolization,” RADIOLOGY, vol. 235, no. 2, pp. 659-667, May 2005; Wu, et al., “Feasibility of US-guided high-intensity focused ultrasound treatment in patients with advanced pancreatic cancer: Initial experience,” RADIOLOGY, vol. 236, no. 3, pp. 1034-1040, September 2005; Yuh, et al., “Delivery of systemic chemotherapeutic agent to tumors by using focused ultrasound: Study in a murine model,” RADIOLOGY, vol. 234, no. 2, pp. 431-437, February 2005; Blana, et al., “First analysis of the long-term results with transrectal HIFU in patients with localized prostate cancer,” EURO UROLOGY, vol. 53, no. 6, pp. 1194-1203, June 2008; Uchida, et al., “Transrectal high-intensity focused ultrasound for the treatment of localized prostate cancer: Eightyear experience,” Int. J. UROLOGY, vol. 16, no. 11, pp. 881-886, November 2009; and Hindley, et al., “MRI guidance of focused ultrasound therapy of uterine fibroids: Early results,” Am. J. ROENTGENOLOGY, vol. 183, no. 6, pp. 1713-1719, December 2004).
HIFU offers some unique advantages as a form of non-ionizing radiation suitable for the localized treatment of deep-seated tumors in a noninvasive or minimally invasive manner (see, Sanghvi, et al., “New developments in therapeutic ultrasound,” IEEE Eng. Med. Biol. Mag., vol. 15, no. 6, pp. 83-92, November/December 1996). Image guidance using diagnostic MRI and ultrasound (see, Tempany, et al., “MR imaging-guided focused ultrasound surgery of uterine leiomyomas: A feasibility study,” Radiology, vol. 226, pp. 897-905, November 2003; and Sanghvi, et al., “Noninvasive surgery of prostate tissue by high-intensity focused ultrasound,” IEEE Trans. Ultrason., Ferroelectr., Freq. Contr., vol. 43, no. 6, pp. 1099-1110, November 1996) has led to increased acceptance of HIFU as a noninvasive therapeutic tool. Currently, HIFU is approved worldwide for use in the treatment of uterine leimyomas and prostate cancer. The HIFU beam experiences minimum distortion when focusing at the target sites by utilizing a noninvasive probe for the treatment of the uterine leimyomas and an intracavitary transducer for the prostate (see, Chan, et al., “An imageguided high intensity focused ultrasound device for uterine fibroids treatment,” Med. Phys., vol. 29, pp. 2611-2620, 2002; and Poissonnier, et al., “Control of prostate cancer by transrectal HIFU in 227 patients,” Eur. Urol., vol. 51, pp. 381-387, 2007).
During the treatment session, image guidance is vital to target the treatment location and to avoid the potential for collateral damage to the intervening tissue in the path of the HIFU beam. Temperature-sensitive MRI has been used in monitoring the application of HIFU in the treatment of uterine fibroids, and ultrasound has been shown to provide adequate feedback in guiding the HIFU treatment of prostate cancer.
Current clinical HIFU systems employ concave mechanically scanned transducers with relatively low fnumber (i.e., to provide high focusing gain). Both single-element and (coarsely sampled) array transducers are currently being used. Array transducers for generating HIFU beams offer additional advantages of compensating for tissue heterogeneities in the path of the HIFU beam (see, Chapelon, et al., “New piezoelectric transducers for therapeutic ultrasound,” Ultrasound Med. Biol., vol. 26, pp. 153-159, 2000; Pernot, et al., “High power density prototype for high precision transcranial therapy,” in Proc. 3rd Int. Symp. Ther. Ultrasound, 2003, vol. 1, pp. 405-410; Hynynen et al., “Trans-skull ultrasound therapy: The feasibility of using image-derived skull thickness information to correct the phase distortion,” IEEE Trans. Ultrason., Ferroelectr., Freq. Control, vol. 46, no. 5, pp. 752-755, May 1999; Ishida, et al., “Development and animal experiment of variable focusing HIFU system for prostate cancer treatment,” in Proc. 3rd Int. Symp. Ther. Ultrasound, 2003, vol. 1, pp. 382-387; Seip, et al., “High-intensity focused ultrasound (HIFU) phased arrays: Recent developments in transrectal transducers and driving electronics,” in Proc. 3rd Int. Symp. Ther. Ultrasound, 2003, vol. 1, pp. 423-428; Curiel, et al., “1.5-D high intensity focused ultrasound array for non-invasive prostate cancer surgery,” IEEE Trans. Ultrason., Ferroelectr., Freq. Control, vol. 49, no. 2, pp. 231-242, February 2002; Martin, et al., “Investigation of HEM produced emulsion for acoustic hemostasis,” in Proc. 3rd Int. Symp. Ther. Ultrasound, 2003, vol. 1, pp. 351-356; and Aubry, et al., “Transcostal high-intensity-focused ultrasound: Ex vivo adaptive focusing feasibility study,” Phys. Med. Biol., vol. 53, pp. 2937-2951, 2008).
Depending on the size and distribution of the array elements, amplitude and/or phase compensation of the driving signals to the elements can be used to refocus the HIFU beam at the target in the presence of tissue aberrations. This, of course, assumes that information about tissue aberration is reliably measured or estimated. One way to estimate these aberrations is by using 3-D numerical modeling of the acoustic wave propagation based on tissue parameters from pretreatment X-ray computed tomography (CT) or MRI patient datasets (see, Tanter, et al., “Focusing and steering through absorbing and aberrating layers: Application to ultrasonic propagation through the skull,” J. Acoust. Soc. Amer., vol. 103, pp. 2403-2410, 1998; and Sun, et al., “Focusing of therapeutic ultrasound through a human skull: A numerical study,” J. Acoust. Soc. Amer., vol. 104, pp. 1705-1715, 1998). This approach has been suggested for focusing HIFU beams through the skull, but it is only of limited value when targeting tumors in abdominal organs where motion is significant. Alternatively, implantable hydrophones can be used to measure the array directivity at or near the target and refocus the beam based on phase-conjugation or time-reversal methods (see, Seip, et al., “Dynamic focusing in ultrasound hyperthermia treatments using implantable hydrophone arrays,” IEEE Trans. Ultrason., Ferroelectr., Freq. Control, vol. 41, no. 5, pp. 706-713, September 1994). This approach was suggested for focusing hyperthermia arrays where the acoustic sensors can be integrated with the necessary temperature sensors.
Recently the concept of dual-mode ultrasound array (DMUA) systems for image-guided application of therapeutic HIFU have been discussed (see, Ebbini, et al., “Lesion formation and visualization using dual-mode ultrasound phased arrays,” in Proc. IEEE Ultrason. Symp., October 2001, vol. 2, pp. 1351-1354; Steidl, et al., “Dual-mode ultrasound phased arrays for noninvasive surgery: Post-beamforming image compounding algorithms for enhanced visualization of thermal lesions,” in Proc. IEEE Int. Symp. Biomed. Imag., July 2002, pp. 429-432; Yao and Ebbini, “Real-time monitoring of the transients of HIFU-induced lesions,” in Proc. IEEE Ultrason. Symp., October 2003, vol. 1, pp. 1006-1009; Yao and Ebbini, “Dual-mode ultrasound phased arrays for imaging and therapy,” in Proc. IEEE Int. Symp. Biomed. Imag., April 2004, vol. 1, pp. 25-28; and Ebbini, et al., “Dual-mode ultrasound phased arrays for image-guided surgery,” Ultrason. Imag., vol. 28, pp. 201-220, 2006). The advent of piezo-composite transducer technology has provided transducers capable of producing high-power levels suitable for therapy with reasonably wide bandwidth suitable for imaging (see, Fleury, et al., “New piezocomposite transducers for therapeutic ultrasound,” in Proc. 2nd Int. Symp. Ther. Ultrasound, 2002, vol. 1, pp. 428-436). Furthermore, piezo-composite technology results in array elements with low lateral cross-coupling leading to more predictable element and beam patterns, both in imaging and therapy modes. A number of approaches for improving the image quality of a prototype DMUA that was originally optimized for therapeutic performance have been investigated (see, Ebbini, “Deep localized hyperthermia with ultrasound phased arrays using the pseudoinverse pattern synthesis method,” Ph.D. dissertation, Univ. Illinois, Urbana, 1990; and Wan and Ebbini, “Imaging with concave large-aperture therapeutic ultrasound arrays using conventional synthetic-aperture beamforming,” IEEE Trans. Ultrason., Ferroelectr., Freq. Control, vol. 55, no. 8, pp. 1705-1718, August 2008), including conventional imaging in sector scan format (see, Simon, et al., “Combined ultrasound image guidance and therapy using a therapeutic phased array,” SPIE Med. Imag., vol. 3341, pp. 89-98, May 1998); and Cartesian coordinates using synthetic aperture (SA) and single transmit focus (STF) imaging (see, Ebbini, et al., “Lesion formation and visualization using dual-mode ultrasound phased arrays,” in Proc. IEEE Ultrason. Symp., October 2001, vol. 2, pp. 1351-1354), harmonic and nonlinear quadratic imaging (see, Yao, et al., “Enhanced lesion visualization in image-guided noninvasive surgery with ultrasound phased arrays,” in Proc. 23rd Annu. Int. Conf. IEEE Eng, Med. Biol. Soc., October 2001, vol. 3, pp. 2492-2495, nonlinear frequency compounding (see, Steidl, et al., “Dual-mode ultrasound phased arrays for noninvasive surgery: Post-beamforming image compounding algorithms for enhanced visualization of thermal lesions,” in Proc. IEEE Int. Symp. Biomed. Imag., July 2002, pp. 429-432), and the use of coded excitation with pseudoinverse filtering to balance axial and lateral resolution (see, Wan and Ebbini, “Imaging with concave large-aperture therapeutic ultrasound arrays using conventional synthetic-aperture beamforming,” IEEE Trans. Ultrason., Ferroelectr., Freq. Control, vol. 55, no. 8, pp. 1705-1718, August 2008; and Shen and Ebbini, “Filter-based coded-excitation system for high-speed ultrasonic imaging,” IEEE Trans. Med. Imag., vol. 17, no. 6, pp. 923-934, December 1998).
In addition to problems associated with the possible obstruction of the HIFU beam by structures resulting in inadequate therapy at the target and/or treatment-limiting pain or damage to normal tissues in the path of the beam (e.g. ribs when targeting liver tumors), HIFU suffers from other limitations that may hinder a wider acceptance of this modality. For example, one limitation is the long treatment time compared to competing minimally invasive modalities. For example, a tumor may be treated in 15 minutes using RF ablation, but may require 2-3 hours using a conventional HIFU protocol (e.g., raster scan of small ablations within the focal region of the HIFU application).
Various technologies for noninvasive application of therapeutic HIFU have been discussed. For, example, such technologies may include piezo-composite array transducer technology (see, Chapelon, et al., “New piezoelectric transducers for therapeutic ultrasound,” ULTRASOUND IN MEDICINE AND BIOLOGY, vol. 26, no. 1, pp. 153-159, January 2000) and noninvasive thermometry (see, Seip and Ebbini, “Non-invasive estimation of tissue temperature response to heating fields using diagnostic ultrasound,” IEEE Trans. Biomed. Eng., vol. 42, no. 8, pp. 828-839, 1995; Seip, et al., “Noninvasive real-time multipoint temperature control for ultrasound phased array treatments,” IEEE TRANSACTIONS ON ULTRASONICS FERROELECTRICS AND FREQUENCY CONTROL, vol. 43, no. 6, pp. 1063-1073, November 1996; Simon, et al., “Two-dimensional temperature estimation using diagnostic ultrasound,” IEEE Trans. Ultrason., Ferroelect., Freq. Contr., vol. 45, pp. 989-1000, July 1998; Salomir, et al., “Hyperthermia by MR-guided focused ultrasound: Accurate temperature control based on fast MRI and a physical model of local energy deposition and heat conduction,” Magnetic Resonance in Medicine, vol. 43, pp. 342-347, 2000; Vanne and Hynynen, “MRI feedback temperature control for focused ultrasound surgery,” Physics in Medicine and Biology, vol. 48, no. 1, pp. 31, 2003; and Souchon, et al., “Monitoring the formation of thermal lesions with heat-induced echo-strain imaging: A feasibility study,” Ultrasound in Medicineand Biology, vol. 31, pp. 251-259, 2005).
Phased array applicators offer unparalleled level of spatial and temporal control over the heating pattern, including simultaneous heating at multiple-focus locations (see, Ebbini, Deep Localized Hyperthermia with Ultrasound Phased Arrays Using the Psudoinverse Pattern Synthesis Method, Ph.D. thesis, University of Illinois, 1990; and Ebbini and Cain, “Experimental evaluation of a prototype cylindrical section ultrasound hyperthermia phased-array applicator,” IEEE TRANSACTIONS ON ULTRASONICS FERROELECTRICS AND FREQUENCY CONTROL, vol. 38, no. 5, pp. 510-520, September 1991). This has many potential advantages in thermal therapy (see, Ebbini, et al., “Dual-mode ultrasound arrays for image-guided surgery,” Ultrasonic Imaging, vol. 28, pp. 65-82, April 2006).
Temperature imaging using MRI is available on clinical MR-guided HIFU systems (MRgFUS) and can be credited in the increased awareness and acceptance of this form of noninvasive surgery (see, Salomir, et al. (2000); and Vanne and Hynynen (2003)). Feedback control algorithms of HIFU fields based on noninvasive temperature imaging using MRI has been described (see, Salomir, et al. (2000); Smith, et al., “Control system for an MRI compatible intracavitary ultrasound array for thermal treatment of prostate disease,” INTERNATIONAL JOURNAL OF HYPERTHERMIA, vol. 17, no. 3, pp. 271-282, May-June 2001; Mougenot, et al., “Automatic spatial and temporal temperature control for MR-guided focused ultrasound using fast 3D MR thermometry and multispiral trajectory of the focal point,” MAGNETIC RESONANCE IN MEDICINE, vol. 52, no. 5, pp. 1005-1015, November 2004; Sun, et al., “Adaptive real-time closed-loop temperature control for ultrasound hyperthermia using magnetic resonance thermometry,” CONCEPTS IN MAGNETIC RESONANCE PART B-MAGNETIC RESONANCE ENGINEERING, vol. 27B, no. 1, pp. 51-63, October 2005; and Mougenot, et al., “Three-dimensional spatial and temporal temperature control with MR thermometry-guided focused ultrasound (mrghifu),” Magnetic Resonance in Medicine, vol. 61, pp. 603-614, 2009).
Ultrasound temperature estimation has also been described (see, Simon, et al., “Two-dimensional temperature estimation using diagnostic ultrasound,” WEE Trans. Ultrason., Ferroelect., Freq. Contr., vol. 45, pp. 989-1000, July 1998; Miller, et al., “Fundamental limitations of noninvasive temperature imaging by means of ultrasound echo strain estimation,” Ultrasound in Medicine and Biology, vol. 28, pp. 1319-1333, 2002; and Pernot, et al., “Temperature estimation using ultrasonic spatial compounding,” hEE Trans. Ultrason., Ferroelect., Freq. Contr., vol. 51, no. 5, pp. 606-615, 2004), as well as a photoacoustic-based approach (see, Pramanik and Wang, “Thermoacoustic and photoacoustic sensing of temperature,” JOURNAL OF BIOMEDICAL OPTICS, vol. 14, no. 5, September-October 2009). A number of different ultrasound thermography methods have been proposed (see, Seip and Ebbini, “Non-invasive estimation of tissue temperature response to heating fields using diagnostic ultrasound,” IEEE Trans. Biomed. Eng., vol. 42, no. 8, pp. 828-839, 1995; Maass-Moreno and Damianou, “Noninvasive temperature estimation in tissue via ultrasound echo shifts. Part I. Theoretical model,” The Journal of the Acoustical Society of America, vol. 100, pp. 2514-2521, 1996; and Arthur, et al., “In vivo change in ultrasonic backscattered energy with temperature in motion-compensated images,” INTERNATIONAL JOURNAL OF HYPERTHERMIA, vol. 24, no. 5, pp. 389-398, 2008).